Cancer therapy utilizing external high energy x-ray, RT (Radiation Therapy), has developed extensively since the introduction of IMRT (Intensity Modulated Radiation Therapy). Traditional RT utilizes radiation beams with homogenous dose distribution in the primary beam. However, due to variations in beam attenuation across the field caused by irregular shape of the patient anatomy and variation in tissue composition, the actual dose distribution in the patient tumor is often more or less heterogenic. In IMRT the dose distribution is optimized to be homogenous in the target. This is accomplished by calculating and delivering the primary beam as a non-flat, intensity modulated beam. The tool used to plan and optimize the treatment is a TPS (Treatment Planning System).
To create the intensity modulated beam, the primary beam is shielded during different time periods for different areas in the field to be treated. This is normally performed by using a MLC (Multi Leaf Collimator) where each individual collimator leaf is controlled separately. A typical MLC have 60-80 pairs of 5 mm leafs, wherein each pair is capable of opening a radiation slit of up to 40 cm.
In the first generation of method for intensity modulated treatment, the treatment delivery was provided in a fixed number of projections. In each projection, either a number of fixed fields with different shapes where provided to create the Intensity Modulated field, so called “step and shoot”, or, alternatively, the MLC-leafs were moved when the beam was ON, so called “sliding window”.
The introduction of fixed-projection IMRT improved the treatment results compared to previous fixed field technique but the trade-off was significantly increased treatment time and more time was thereby needed to treat each individual patient, which also increased the costs involved in treatment significantly.
Additionally the creation of the intensity modulated fields increased the treatment complexity and new QC (Quality Control) was required to make sure the treatment was given in accordance with the plan created with the TPS. The QC method PTV, “Pre Treatment Verification”, was therefore introduced. In PTV, the 3D dose distribution in a phantom (patient substitution) is calculated using the TPS tools for a specific plan and for a specific patient. The phantom doesn't have the same shape and heterogeneities as the patient and thereby the dose distribution will not be the same as in the patient. However, the phantom can be irradiated using the patient specific plan and the dose distribution measured inside the phantom can thereby be compared with the planed dose distribution. If the measured dose distribution in the phantom correlates with the dose distribution in the treatment plan for the phantom, it has been proven that the planned treatment can be given as intended.
The ideal phantom would be similar to the patient in shape and density with detectors in the full 3D volume. That is currently not possible due to costs etc. To measure inside the real patient would require a large number of measurement points which is not feasible either. A technique that optimizes the detector configuration in relation to the requirements is described in the U.S. Pat. No. 7,371,007.
Depending on the intended use for such a phantom, the requirement on the isotropy (measurement dependency on incident angle) might vary. If the phantom is to cover from “head and neck” to pelvis, the requirement on isotropy is somewhat limited due to that the incident angle of the beam is limited to solid angle +/−30 degrees in a 360 degree rotation. If the intended use is to cover also full head including brain; the incident beam might be almost 4π, i.e. almost any beam direction. High accuracy measurements will then require even higher demands on the detector system to be isotropic, i.e. directional independent.
An isotropic detector should also fulfill other requirements such as high spatial resolution, energy independency, dose linearity, dose per pulse linearity, low temperature dependency, stability, high signal/noise ratio, radiation tolerances, real time measurement, configurable in arrays of detectors and not least cost effect to make them useful in practice.
In a first approximation, for a detector that is small compared to the range of the secondary electron that creates the dose (energy per mass), most of the dose detected by the detector originate from electrons generated outside the detector itself and enters the active volume and deposit the dose. If the surrounding material of the small detection volume is inhomogeneous, the generation of the secondary electrons will depend on the mass density and electron density of the surrounding materials and if it varies the detector will not be isotropic, thereby it is important to have a homogeneous or symmetric surrounding of the detector volume.
In a traditional diode-detector-chip part of the chip is active and part of it is inactive (bulk) in the collection of free charges. The geometric shape is thereby none uniform around the active part of the chip and the secondary electrons from the surrounding will thereby not be uniform; the creation of the local dose will become directional dependent.
Secondary electrons that enter the active volume via the bulk silicon respectively directly via e.g. water equivalent surrounding the active volume will create different amounts of secondary electrons and thereby the imparted dose will vary with the incident angle.
Attempts to provide isotropic detectors have been made, in the early 1990's a double chip (sandwich) detector was constructed to reduce the effect mentioned above. The sandwich detector was more isotropic than a single chip detector but is however impaired with drawbacks. For example, the construction requires that the active part is symmetrically placed in relation to the bulk-silicon and symmetric in shape which, for example, makes it difficult to manufacture. These geometrical asymmetries may lead to a signal response which is not independent of the direction of the incoming radiation.
Another issue in detectors where the density differs from the surrounding material is energy dependency. The attenuation of low energy (<200 keV) photons is for example up to 7 times higher than in water, which leads to an energy dependency in a sandwich construction being more pronounced than in a single chip detector where almost 50% of the surface is facing the surrounding material.
Another attempt to provide an isotropic detector is presented in US 2009/0057562, where a method, apparatus, and computer program for measuring the dose, dose rate or composition of radiation are disclosed. In one embodiment, an apparatus for detecting and measuring an ambient unknown radiation field includes a large number of detector chips that are facing in different directions is disclosed. However, the detector presented in US 2009/0057562 may also be impaired with problems related to the symmetry of the detector chips, which, in turn, may lead to a signal response which is not independent of the direction of the incoming radiation.
Thus, there is still a need within the art for isotropic detectors having a signal response which is independent of the direction of the incoming radiation that at the same time fulfill other important requirements, such as high spatial resolution, energy independency, dose linearity, dose per pulse linearity, low temperature dependency, stability, high signal/noise ratio, radiation tolerances, real time measurement, on a detector system for use in measuring radiation dose in photon or electron fields such as for radiation medicine, including radiotherapy and radiation based diagnosis.